Ultrasonic imaging is widely used in medicine because it does not subject the subject to ionizing radiation and is considerably less expensive than systems based on magnetic resonance imaging. In one application of ultrasound imaging, a probe containing an ultrasonic transducer is inserted into the body area to be imaged. The transducer transmits an acoustic pulse into the body tissues, and detects the reflections of the pulse at tissue boundaries due to differences in acoustic impedance, as well as the backscattered sound from acoustically heterogeneous tissue. The differing times taken for the transducer to receive the reflected and backscattered ultrasound correspond to the differing distances of the tissues from the transducer. By stepping or sweeping the transducer through a set of selected angles, a two-dimensional ultrasound image corresponding to a map of the acoustic impedance boundaries and backscattering coefficients can be obtained. From this image, the condition of the body tissues can be determined. For example, the method of intravascular ultrasound (IVUS) sequentially transmits ultrasound pulses in equally spaced increments around all or part of a circle to obtain cross-sectional images of blood vessels, most often coronary arteries demonstrating areas of atherosclerotic plaque, calcification, etc.
Generally, there are two types of ultrasonic probes for IVUS imaging. The first type employs a synthetic aperture technique. For example, U.S. Pat. No. 4,917,097 (Proudian, et al.) and U.S. Pat. No. 5,186,177 (O'Donnell, et al.) teach probes based on synthetic apertures. Generally, these probes utilize the sequential excitation of selected elements in an array of transducer elements to generate a sound pulse traveling in a particular direction which is determined by the elements excited and the relative phases of the excitation signals applied thereto.
The second type of IVUS probe scans the tissue by mechanical rotation of a transducer that emits pulses in a predetermined direction. The mechanically rotated type includes a few subclasses. In the first subclass, either a distal (remote from the operator) transducer or a mirror is rotated from the proximal end of the catheter by an extended drive shaft with a proximal motor (U.S. Pat. Nos. 4,794,931 and U.S. Pat. No. 5,000,185 (Yock)). In the second subclass, the rotation is confined to the distal end, where either a miniature motor (U.S. Pat. No. 5,240,003 (Lancee, et al.) and U.S. Pat. No. 5,176,141 (Bom, et al.)) or a fluid driven turbine is used to rotate the transducer or the mirror (U.S. Pat. No. 5,271,402 (Yeung and Dias)). In a third subclass, a stationary proximal transducer is acoustically coupled to a rotating acoustic waveguide that conducts the sound to the distal end (U.S. Pat. No. 5,284,148 (Dias and Melton)). In a fourth subclass (U.S. Pat. No. 5,509,418 (Lum, et al.)), a turbine is rotated by an acoustic signal generated outside the vessel to direct another ultrasonic signal in a rotating fashion. In the final subclass (U.S. Pat. No. 5,507,294 (Lum, et al.)), an external driving member rotates a tube to rotate a reflecting element at the tip of the tube to reflect ultrasound.
Presently, probes that direct ultrasonic pulses by mechanical rotation are more widely used than probes that electronically aim the pulses. The mechanical approach can be implemented using a single transducer, while the electronic approach requires an array of transducers to be contained in the distal end. Accordingly, the array requires a larger catheter. In many applications such as imaging blood vessels, minimizing the size of the catheter is essential.
Mechanical probes, however, can introduce distortions into the images resulting from uncertainties in the speed of rotation of the transducer. Typically, the distal end of the probe is assumed to rotate at a constant speed and ultrasound pulses are transmitted at regular time intervals. The angular position of each scan line is assumed to change by the same amount between each successive pulse. If, however, the angular rotation is not uniform, the angular positions of the scan lines corresponding to the various pulses will be in error and the resulting image will be distorted.
One cause of non-uniform angular velocity in the type of catheter that uses a driveshaft is the existence of mechanical friction between the spinning driveshaft and the surrounding stationary sheath. The catheter must bend numerous times through the tortuous path involved in placing the distal end at the desired location by traversing a blood vessel. Although the proximal end of the catheter is rotating at the desired angular velocity, any binding of the catheter along its length will lead to a distal angular velocity that is different from the desired velocity at various points of the full circle. The average velocity will be the same at the proximal and distal ends, and thus the distal end will sometimes be rotating too quickly, and sometimes too slowly. In general, the error is observed to be substantially the same on subsequent revolutions of the catheter. Thus, the image generated appears to be distorted even when large numbers of measurements are made.
Techniques for correcting for non-uniform rotation have been proposed. For example, U.S. Pat. No. 5,485,845 (Verdonk, et al.) describes a technique for detecting the non-uniform angular velocity of IVUS transducers by using an array of beacons positioned on the sheath. This approach determines the average angular velocity of the transducer between each pair of beacons. The accuracy of this approach is determined by the number of beacons and the placement thereof. To obtain high accuracy, large numbers of beacons are needed. However, the beacons can interfere with the image since the beacons form "bright" spots and shadows on the image. This interference increases with the number of beacons.
U.S. Pat. No. 5,699,806 (Webb, et al.) describes a technique for calculating the non-uniformity from the distorted images. This technique depends on the statistical distribution of the speckle pattern observed in ultrasound images of backscattering tissue. This pattern is absent in some imaging situations. For example, at the branch point between two blood vessels, there may be a region in which no reflected signal is obtained. Similarly, if the blood vessel has a calcified region, the strong reflection from the calcified region may interfere with the reflections and scattering generated by the tissue behind the calcified region, and hence, the detection of the angular velocity of the transducer.
Broadly, it is the object of the present invention to provide an improved ultrasound catheter and calibration method.
It is a further object of the present invention to provide an ultrasound catheter whose angular position can be determined without obstructing portions of the image.
It is a still further object of the present invention to provide an ultrasound imaging system that does not depend on the statistical properties of the image to determine the angular position of the transducer.
These and other objects of the present invention will become apparent to those skilled in the art from the following detailed description of the invention and the accompanying drawings.